Magnetic Resonance Imaging (MRI) is now a widely accepted and commercially available technique for obtaining digitized visual images representing the internal structure of objects (such as the human body) having substantial populations of nuclei which are susceptible to NMR phenomena. In MRI certain nuclei of a body to be imaged are polarized by imposing a strong main magnetic field H.sub.0 on them. Then, selected nuclei are excited by imposing on the nuclei an RF signal at a particular NMR frequency. By spatially distributing the localized magnetic fields, imposing an RF signal to them, and then suitably analyzing the resulting RF responses from the nuclei, a map or image of relative NMR responses as a function of the location of the nuclei can be determined. Following a Fourier analysis, data representing the population of nuclei in space can be displayed on a CRT.
As shown in FIG. 1, the NMR imaging system typically includes a magnet 10 to impose the static magnetic field, gradient coils 12 for imposing spatially distributed magnetic fields in the three orthogonal coordinates, and RF coils 14 to transmit and receive RF signals to and from the selected nuclei. The NMR signal received by the coil 14 is transmitted to a computer 16 which processes the data into an image displayed on the display 18. The computer 16 also controls the operation of the RF coils 14 and gradient coils 12 through the RF amplifier 20 and gradient amplifiers 22, respectively.
An example operation of how the various coils produce the NMR signal in 2D MRI is shown in FIG. 2, which is a graphical representation of a typical NMR acquisition sequence at steady-state. First, a gradient field G.sub.Z is applied to sensitize a slice of nuclei in the body to be imaged to a particular RF resonance frequency. An RF nutation pulse is then applied at the particular frequency to force some of the nuclei to precess in perpendicular to the main field. Thereafter, a changing gradient field G.sub.Y phase encodes the nuclei in the Y-axis direction by altering the frequency differences (and hence the phase) between nuclei in different locations along the Y-axis. A gradient field G.sub.X first dephases and then rephases the nuclei to form a field-echo NMR signal. During the field-echo, the gradient field G.sub.X frequency encodes the selected slice of nuclei in the X-axis direction. The resultant NMR signal is then read and analyzed by Fourier analysis. The frequency (domain) plot of that analysis is then scaled to render information about the nuclei population in Fourier space, which corresponds to an X-Y-Z position.
The above examples rely on a non-varying main magnetic field that polarizes the nuclei. Otherwise, time-variation of the main magnetic field can cause anomalies in the slice selection, the phase encoding, and the frequency encoding steps during data acquisition. Ideally, the main magnetic field is assumed to be perfectly stable. The phase encoding gradient field, G.sub.Y, for example, would then superimpose a magnetic field varying linearly in the Y-direction while being constant in all other dimensions. The total magnetic field in the Y-direction in this example would be the sum of H.sub.0 and G.sub.Y. The phase of the nuclei is then locally encoded only by their position along the Y-direction.
In reality, however, small spurious changes occur in the main magnetic field during MRI data measurement processes. The changes over time of the main magnetic field can introduce spurious changes in slice position, and in the phase and frequency of the nuclei. Unless these spurious changes are corrected, the original spatial encoding will be aberrant, causing artifacts in the resultant NMR image.
For example, the magnetic gradient fields produce eddy currents, which induce a time-varying magnetic field, thereby causing variations in the main magnetic field. This is particularly true of the magnetic gradient pulses used to achieve slice/volume selective NMR responses, which are of substantially greater intensity/duration than other magnetic gradient pulses and, accordingly, are often the principal source of spurious magnetic fields due to induced eddy currents. Other system or environmental characteristics (such as temperature, system vibration, etc.) can also cause variations and drift in the main field.
This drift can have an impact on image accuracy. In Fourier Transform (FT) MRI, the acquisition time of each echo signal (the "read-out time") is on the order of milliseconds, while the total data collection time (the "total time") is on the order of minutes. When the main magnetic field changes on a time scale which is slow for the "read-out time" but fast for the "total time", image artifacts will be generated, mainly in the phase-encoding direction(s). These artifacts, in many cases, render the images useless. That is, the field variations induce phase errors in the recorded signals, causing the Fourier analysis to misidentify the spatial distribution of nuclei populations. The misidentifications appear as artifacts on the CRT display.
One spectroscopic method previously developed (U.S. Pat. No. 4,885,542) for correcting spurious changes in the main magnetic field requires high-resolution NMR signals be acquired and used to quantify the time course of the main field. These signals record changes in the main field which are then used during data processing to correct for the field change and to eliminate the associated image artifacts. In particular, according to this method, a template or calibration NMR response is obtained from each slice without phase or frequency encoding. The strength of the main field is quantified from the frequency of the NMR signal. The MRI signals for the slice are corrected in accordance with the calibration response. In practice, for example, if one desired 128 projections (phase-encoding steps), 137 measurement cycles may be employed with the extra 9 cycles being calibration cycles for recording the main field and for correcting the remaining 128 cycles used in the image reconstruction.
Although the spectroscopic approach produces, in many cases, satisfactory results, it has limitations. First, it requires acquisition of high-resolution NMR signals for calibration. This unfortunately demands a long acquisition duration to achieve sufficient frequency resolution, meaning that the sequence segment for acquiring the field information has very different RF and field gradient characteristics than the imaging segment. Insertion of such a segment amid the image data acquisition can significantly alter the image contrast. Second, when the imaging sequence requires steady-states of the nuclear spin precession, insertion of the spectroscopic calibration segment will interrupt the steady-states. Such perturbation to the imaging sequence will not only affect image contrast, it will produce image artifacts by itself.